Conventional computed tomography (CT) systems with an X-ray tube operating at a constant voltage and a constant current face image-quality problems in scanning various parts of a patient in one acquisition or scanning one part where the attenuation strongly depends on a view angle. Since the physical dimensions of certain body parts significantly vary with respect to views, the transmitted X-ray attenuates at a different level. In other words, since a prior art X-ray energy source provides the same spectrum and flux of the X-ray beams to acquire projection data sets regardless of the physical dimensions of a scanned part, an image quality varies depending upon the physical dimension and other variables such as a material distribution. To improve image quality, as the voltage and current are fixedly increased with respect to large dimensional views, a patient is exposed to an unnecessary amount of high radiation while projection data are acquired from small dimensional views. On the other hand, as the voltage and current are fixedly decreased with respect to small dimensional views, image quality suffers from artifacts due to the poor quality of data from large dimensional views. The constant exposure level prevents the optimal data acquisition. Consequently, the patient may be either overexposed or underexposed during the data acquisition.
With respect to the above issue, prior art techniques modulated either a tube voltage or a tube current of a single energy X-ray source. For example, Japanese Patent Publication 53-110495 discloses that an amount of X ray at a detector is kept constant across various parts of the scanned body by modulating the tube voltage based upon a feedback signal from the X-ray detector. In other words, assuming the same X-ray attenuation coefficient through various parts of a different thickness, a large amount of high-energy X ray is projected into a thicker portion while a small amount of low-energy X ray is projected into a thinner portion so that the attenuated x-radiation is constant at the detector across the various portions via a feedback control.
The above prior art technique assumes the constant attenuation coefficient across various parts of the patient body. In contrast to the assumption, various body parts have a different attenuation profile. Furthermore, the voltage is modulated while the scanning process is taking place. In other words, since the voltage is being modulated at various angles about the scanning axis and also at various positions along the same scanning axis, the acquired projection data are not compatible with each other for the image reconstruction processes. For example, for a given projection, there is only a single set of projection data at a particular voltage level, and this projection data cannot be generally combined with another set of projection data at a different energy level due to the energy-dependent nature of the attenuation coefficient. Consequently, image reconstruction fails to yield a desirably artifact-free image.
In this regard, Japanese Patent 2704084 discloses that an attenuated X-ray level is kept substantially constant at various angles and positions along the scanning direction by modulating a tube current level. In order to accomplish the substantial consistency, an attenuated x-radiation value at the present angle and position is approximated to an average attenuated x-radiation value at the previously detected angle and position based upon current modulation. In other words, the tube voltage is kept constant throughout the scanning process while the tube current alone is modulated to avoid the above described projection data incompatibility.
The current modulation is generally useful in controlling the noise level but not the penetration of x-ray beam that relates to the dose efficiency. Roughly speaking, the higher a tube current level is during the projection data acquisition, the lower a noise level becomes in the acquired projection data sets. More precisely, the noise level, variance of the projection data, is inversely proportional to the current level. On the other hand, the tube current is proportional to the dosage to the patient. In this regard, the current modulation may improve image quality and achieve an optimal dosage level in the case of fixed dose efficiency. However, the optimal dosage level can be further reduced by improving the dose efficiency through adjusting the penetration of X-ray beam.
In order to improve the patient safety, an optimal dose efficiency level must be achieved during the CT scanning procedure. That is, the patient is exposed to a minimally necessary radiation level while the image quality in the reconstructed image is not sacrificed. The dose efficiency is generally defined to be (S/N)2 divided by a dose level, where (S/N) is a signal-to-noise ratio. The dose efficiency depends on x-ray spectrum, detector characteristics and attenuation of the imaged subject. Thus, an optimal dose efficiency level is obtained for a particular material of a certain physical size by controlling the spectrum through the tube voltage modulation.
From the above background information, it appears necessary to modulate the tube voltage level in order to optimize the dose efficiency level since a current modulation alone does not achieve the ultimate goals both in patient safety and image quality. On the other hand, from the prior art attempts, the voltage modulation has an inherent issue of data incompatibility among the projection data if images are reconstructed with a single energy CT reconstructor. In this regard, more recent prior art techniques involve dual energy imaging techniques.
Dual energy imaging in CT has been a promising technique since the first days of CT and was even mentioned in Godfrey Hounsfield's paper (1973) that introduced CT. The basic idea is to acquire two data sets at low and high energy levels and to use the pairs of the data sets to deduce additional information about the patient.
The physical basis of dual energy imaging includes two main mechanisms of the interaction of X rays with matter in the clinically relevant diagnostic energy-range from 30 keV to 140 keV, and the two interactions are photoelectric absorption and Compton scattering, each having its own functional dependence on X-ray energy. Photoelectric absorption is a rapidly decreasing function of energy while Compton scatter is a gentle function of energy. As shown in FIG. 1, the photoelectric interaction is a strong function of the effective atomic number (Z) of the absorbing tissue while scattering is nearly independent of Z. The physics enabled Alvarez and Macovski (1976) to develop a mathematical scheme, called dual-energy decomposition, to use the dual energy information.
In addition to the energy dependence, dual-energy decomposition must take X-ray sources into account. Since commercial clinical CT-scanners generally use polychromatic sources, the mathematics of dual energy imaging is not trivial. In this regard, single-energy imaging with a polychromatic source does not have an exact and analytic solution. One mathematical approach in dual-energy decomposition using a polychromatic source has been described in a related U.S. application Ser. No. 12/361,280 filed on Jan. 28, 2009 and Ser. No 12/106,907 filed on Apr. 21, 2008 as well as in a reference entitled as “Analysis of Fast kV-switching in Dual Energy CT using a Pre-reconstruction Decomposition Technique,” by Yu Zou and Michael D. Silver (2009). In dual energy computed tomography (CT), fast kV-switching techniques generally alternate voltages between projections (also called views) so that the odd (or even) projections correspond to the low (or high) tube voltage. These references are incorporated into the current application by external reference to supplement the specification. Instead of the polynomial approximation method, in the previously proposed approach combining a linear term with a non-linear beam hardening term, an iterative solution to the dual energy data domain decomposition converges rapidly due to the dominant linear term.
In the past two years, prior art attempts have implemented certain dual energy CT systems. For example, Siemens has installed a number of dual source CT-scanners, which is equipped with two X-ray sources, and each runs at a different energy level for generating the two data sets. Another example is that Philips at their Haifa research facility has developed a sandwich detector where the upper layer records the low energy data and the lower layer records the high energy data. A prototype system is installed at the Hadassah Jerusalem Hospital. In this regard, GE has developed a specialized detector using garnet for capture 2496 total projections per rotation (TPPR) at a high speed. The fast detector has been combined with a fast kV-switching X-ray source to acquire the low and high energy data sets.
TABLE 1 below summarizes advantages and disadvantages of selected ways to acquire dual energy data sets. Fast kV-switching techniques change voltages between projections (also called views) so that the odd and even projections respectively correspond to the low or high tube voltage. Among these prior art approaches, the fast kV-switching appears an attractive technique for dual energy acquisition for a number of reasons. Since the dual source CT-scanners and the sandwich detector CT-scanners respectively require additional costs for the dual X-ray sources and the sandwich detectors, they may not be cost-effective to obtain dual energy data sets. Similarly, although GE's detector for fast kV-switching energy CT is not summarized in TABLE 1, the semi-precious gem detector also incurs additional costs. In addition, both the dual source CT-scanners and the sandwich detector CT-scanners must resolve other technical difficulties that are associated with these systems as listed in the table below. On the other hand, although the slow kV-switching does not require additional parts or equipment, dual energy data sets result in poor temporal registration that is off by at least one rotation period as well as poor spatial registration in particular from helical scans. For these reasons, the prior art technologies remain to find a cost effective system and method to utilized the dual energy data for CT.
TABLE 1OptionsAdvantagesDisadvantagesFastTemporal and spatialLimited energy separationkV-registration very good.unless square-wave waveformswitchingData domain methodsdeveloped.(alternatingpossible leading to betterDifficult to equalize dose/noiseviews)IQ and flexibility.between high/low data sets.Helical acquisition noDevelopment time and cost forproblem.fast, switching HVPS.SlowGood energy separation.Poor temporal registration; offkV-Easy to equalizeby at least one rotation period.switchingdose/noise betweenPoor spatial registration,(alternatinghigh/low data sets.especially if doing helical scansrotations)Little equipmentand thus limited to imagedevelopment necessary.domain methods.Little or no added H/WHelical scans may require lowercosts.pitch and thus more dose.DualGood energy separation.Temporal registration off by ¼sourceEasy to equalizeof the rotation period.dose/noise betweenSpatial registration requireshigh/low data sets.tube alignment.Cost of two imaging chains.Field-of-view for dual energylimited by the smaller of thetwo imaging chains.Cross-scatter contamination.SandwichPerfect temporal andLimited energy separation.detectorspatial registration.Cost and development of theData domaindetector.decomposition methodsvalid.Helical acquisition noproblem.
As already shown in TABLE 1, prior art fast kV-switching techniques without the use of dual sources or special detectors nonetheless have both advantages and disadvantages in acquiring dual energy data sets. The prior art fast kV-switching techniques have very good temporal and spatial registrations between corresponding high and low energy projections, which make data domain methods possible and lead to better IQ and flexibility. In addition, prior art fast kV-switching techniques acquire good dual energy data sets also through helical projections. A disadvantage is the one view misregistration between corresponding high and low energy projections. Another problem is the difficulty of high noise in the low energy data because it may be technically difficult to swing the mA as fast as the kV.
Regardless of the clinical significance, several hurdles remain for successful dual energy imaging. One important image quality issue is related to the different dose and noise levels between the two data sets. Depending on how the dual energy is achieved, the low energy data set could be very noisy compared with the high energy data set because X-ray tubes are less efficient at lower voltages, and the lower energy X rays usually have worse penetration in tissues, which will be a problem for larger patients. The same issue may be also problematic with scanning various parts of the same patients since the physical dimensions of these body parts significantly vary.
With respect to dual energy CT, one algorithm is described to modulate a current level in “Dual energy exposure control (DEEC) for computed tomography: Algorithm and simulation study,” Phillip Stenner and Marc Kachelrie, Med. Phys. 35 (11), November 2008. The prior art technique claims that DEEC minimizes the noise in the final monochromatic image while it keeps the dose constant. Alternatively, the prior art technique claims that DEEC minimizes the dose while it keeps the noise constant. In other words, either the dose or the noise is improved by the current modulation in the above DEEC. In this regard, the prior art DEEC still fails to improve the dose efficiency.
Despite the above described prior art techniques, patient safety from X-ray overdose remains to be improved without sacrificing image quality. The advantages of dual energy CT includes some improved image quality and other potentially significant contributions.